The therapeutic effects of heat also stimulate the immunological system. Destructive effects are further observed on the microvasculature of the tumor, which further enhances the heating of poorly vascularized tumor cores. When used in combined protocols with radiotherapy, synergistic effects arise due to a specific effect of heat on hypoxic cells, which enhances the killing effect of ionizing radiation. A further beneficial effect of heat in tumors is the enhancement of the effects of tumoricide drugs in combined treatments with chemotherapy (G. Hahn, Hyperthermia and Cancer, Plenum Press, N.Y.).
A further beneficial use of heat is found in palliative treatments of superficial musculature, subcutaneous tissues and/or joints, with heat usually administered as an adjunct to other therapeutical treatments (J. Lehmann, Therapeutic Heat and Cold, Williams and Wilkins, N.Y., 1982). Moreover, alleviation of prostatism occurs with heating the prostatic gland enlarged by benign prostatic hyperplasia. The beneficial effects of drugs is enhanced in a number of non-tumoral pathologies.
When the heating fields are generated by external applicators, it is often difficult to raise tissue temperatures of circumscribed target volume at depth to therapeutic levels without damaging the access tissues represented by the skin and subcutaneous fat layer. So far, the development of an inexpensive and versatile electromagnetic (EM) applicator to safely deliver localized heat to volumes of variable shape and size underneath the subcutaneous fat layer has presented an insolvable problem.
A hyperthermia EM applicator for subcutaneous and deeper tissues is characterized by the following main features. The penetration depth in a muscle tissue is defined as the depth at which the delivered EM energy, measured in Watt/kg in terms of specific absorption rate (SAR), is attenuated to 50% of the normalizing value measured at one cm depth in the same muscle tissue. The SAR space distribution is calculated in terms of the local E(x,y,z) field and conductivity (x,y,z) distributions as: SAR=.alpha.E.sup.2 /2. Thus, for any applicator, an effective field size (EFS) defined by the 50 percent iso-SAR contour measured at one cm muscle depth, hereinafter referred to as either iso-SAR or EFS, may be used to assess the quality of the heating field of an applicator. The penetration depth for an efficient paracorporal applicator designed for subcutaneous heating should be about 2-3 cm within the muscle tissue in order to guarantee a uniform treatment of most muscle tissues. To this figure, the thickness of the access fat and bone layers have to be added, taking into account that in humans the fat layer varies from a few millimeters to a few centimeters. At the same time, the overheating of these access tissues should not be above an acceptable level.
An ideal applicator should not require a water bolus interposed between the applicator and the body surface. The bolus impedes the location and concentration of heat to a specific area and prevents an applicator from scanning a restricted area of the body surface. In addition, an EFS contour can be easily shaped by a moving applicator to the contours and sizes of the tissue to be exposed or to the anatomy of access site when the applicator is mobile rather than fixed.
Finally, an ideal applicator should exhibit low stray radiation levels and no electric hazards and should be patient customized and safely operated by unspecialized personnel in the clinical environment.
There are some intrinsic limitations in the design of EM applicators that should be considered. It is well known to the experts in electromagnetism that heating with EM sources occurs by two types of coupling between the body tissue and the EM radiator. In a dominant E-field type of coupling, in which E is the electric field component of the EM field, the radiator may be substantially described as a charge source. In a dominant H-field type of coupling, with H being the magnetic field component, the radiator may be substantially described as a current source. The latter coupling is referred to as inductive since the heating is due to the induced E-field and associated induced currents in the conductive and therefore lossy tissue. The main restraint in applicator design as required by the EM theory is that the heating E-field should be substantially directed parallel to the subcutaneous fat-muscle (bone-muscle) interface for minimizing the heating of the highly resistive fat (and bone) layer and improving the safety of the treatment.
Good applicator design requires the locally generated E-field to flow parallel to the body surface. This restraint rules out the use of E-field devices which use high power, direct-contact capacitive electrodes whose E-field lines impinge perpendicularly upon the fat-muscle interface. Such applicators produce unpredictable and delocalized heating field distributions and present treatment safety problems due to the overheating of fatty tissue, even if a cooling bolus is used. H-field coupled devices induce E-fields which comply with the above requirement, and are considered the safest and most practical devices. Such H-field applicators have the added advantage that they require neither direct contact with the body surface nor bolus cooling.
A further restraint is related to the frequency of the EM radiation. External heating with radiation of very short wavelength such as microwaves is minimally effective at depth because the absorption of energy in the access tissues along the heating field pathway is so great that insufficient energy reaches deep-seated tissues. Heating at depth with such short wavelengths entails overheating of the access tissues which would be subject to unsafe high intensity fields. It is well known to the experts in the art of therapeutic heating that the penetration depth of an EM radiator increases in direct proportion to the electric dimension of the radiator. This is a problem in the cases of localized heating of a small target volume for which small apertures have to be used. However, the smaller the aperture, the higher the frequency which lowers the penetration potential. This limitation on the use of higher frequencies comes from EM theory whereby the wavelength of the EM field supported by the modes of a resonant waveguide aperture radiator, hereinafter referred to as a resonant waveguide or resonant aperture radiator, is related to the transverse electric dimensions of the aperture. These in fact define the cutoff frequency of such waveguide, i.e., the lower limit for the working frequency in this modality of propagation. Aperture applicators working at frequencies as low as 27 MHz have been developed. However, in order that such a low frequency will still be above the cutoff frequency, waveguides must be dielectric-filled so that their transverse dimensions are brought down to sizes comparable with those of the tissue to be heated. A water loaded, very heavy and cumbersome waveguide applicator has been proposed (A. Paglione et al., Microwave J., Vol. 24, p. 71, 1981) which exhibits high penetration, but its aperture is far too large for the localized treatment of most subcutaneous tumors. Small size microstrip applicators working at 27 MHz have also been developed (R. H. Johnson et al., Strahlentherapie, Vol. 9, pp. 537-538, 1985) but the penetration is not improved with respect to microwave applicators of comparable aperture size and a thick bolus is required to prevent the strong EM near-fields from overheating access tissues. This results in a substantial reduction of the power density of the field impinging upon the body surface.
Thus, concentrating the heating field in a small cross section while maintaining high penetration are conflicting requirements which have limited the prior art development of resonant aperture applicators. In clinical practice, heating small target volumes by small aperture applicators working in the 200-600 MHz range is feasible with acceptable uniformity of heating for superficial tissues at depths not exceeding -1.5 cm, subcutaneous fat layer included. Within these limits, a full set of dielectric-loaded resonant aperture radiators of varying aperture size and penetration would have to be developed in order to meet wide clinical requirements. In any case, no in-field optimization of these applicators would be possible and precise treatment planning for small size subcutaneous target volumes at substantial depths could not be satisfactorily achieved given the high temperature gradients required for target tissue and the rapidly decaying heating fields with consequent low penetration.
Improvements in penetration and uniformity are obtained by the use of phased arrays in which a multi-element radiator is directed toward the target volume with a multiplicity of coherent electromagnetic heating fields which are controlled in phase, amplitude and orientation to give rise to a positive interference effect when out of phase and thus substantially enhancing the temperature elevation of tissue when in phase. This constitutes a method to focus the heating at a predetermined depth. The constructive interfering superposition of microwave or radio frequency radiation fields have long been employed in hyperthermia. Phased arrays of 4, 8 and even 16 resonant aperture applications are known. These, however, exhibit a complexity of operation and high manufacturing costs which are not rewarded by the small gain in uniformity and penetration obtained.
Fixed aperture applicators developed around undersized, air-filled, below-cutoff waveguides (BCW) have been proposed for hyperthermia therapy (J. Vrba et al., Tesla Electr., Vol. 2, pp. 44-50, 1984; J. Vrba, Czechoslovak Patent 227,270). In the design of these heating devices, the excitation of evanescent modes in the BCW for producing useful heating fields occurs accidentally and in uncontrolled ways and under no theoretical conditions would these devices produce heating fields as versatile and effective as the devices disclosed in the present invention.
Attempts to circumvent the intrinsic limitations of aperture radiators for deep subcutaneous treatments have been made by developing H-field and E-field heating devices working at low frequency and using the well-established technology of the inductive shortwave diathermy (J. Oleson, in IEEE Trans. Biomed. Engin., Vol. BME-31, pp. 91-97, 1984). H-field devices do possess the extremely important feature that in whichever direction the inducing currents flow with respect to the body surface, the locally induced E-field and associated currents will flow parallel to the body surface and to the subcutaneous fat-muscle interface, thus sparing the access fat layer from overheating.
Low frequency H-field devices exploit the quasi-static term of the EM field and appear to be of practical use because (1) they do not require a bolus and (2) are less expensive to manufacture and are of proven technology. These devices are substantially coils of various shapes derived from the flat, spiral or pancake multiturn coil design. They are widely used in shortwave diathermy and are placed externally with their coil plane parallel to the body surface. These coil applicators produce inside the body induced solenoidal E-fields and associated current loops which flow on planes substantially parallel to the plane of the inducing current loops, i.e., to the body surface.
The main limitation in the use of low frequency inductive devices is that the induced current loops exhibit a gradient towards their centroids, where the deposited SAR is vanishing, so that their SAR deposition pattern is non-uniform. Moreover, the penetration potential of multiturn coils is substantially impaired by the presence of large stray E-fields between coil turns, the field lines of which are impinging perpendicularly on the fat-muscle interface, causing subcutaneous fat overheating which limits the power that can safely be used and consequently the penetration depth. For both these limitations, multiturn coil devices are used only occasionally in tumor thermotherapy, where precise and uniform fields are required.
Improvements in penetration have been obtained in producing perpendicular E-field loops inside the body by the use of flexible magnetic flux-guides implemented with toroidal resonators at radio frequency, in which a high density magnetic flux is directed over the body surface by treatment ports suitably oriented (Proc. Hypert. Oncology 1988, T. Sugahara and M. Saito, eds., pp. 829-831, Taylor & Francis, 1989). A substantial part of this improvement is due to the low impedance exhibited by these flux guides, which are similar to a curved solenoid. In fact, the current lines are distributed over the whole toroidal wall and are flowing radially. However, this applicator appears to have a limited capability in determining the heating field size, shape and localization of the induced current loops.
A partial removal of the intra-turn E-fields of multiturn coils has been obtained in single rectangular loop applicators used with their loop plane perpendicular to the body surface, hereinafter referred to as magnetic dipoles or dipoles, which are inducing local E-fields also perpendicular to the body surface. The dipole heating field is characterized by a component due to the loopside proximal to the body surface to which the smaller but out-of-phase field of the distal loopside is superimposed. The negative effect of this out-of-phase field depends on the separation between loopsides, i.e., on the dipole height.
Planar dipoles have also been implemented by a large ribbon-like or sheet conductor working at 150 MHz, including a large metallic backplane. This dipole will be referred to as a distributed current dipole or distributed dipole (J. Bach Andersen et al., IEEE Trans. BME, Vol. 31, pp. 21-27, 1984).
Distributed dipoles at various frequencies have been described by others who have provided them with metal screening boxes which closely wrap the dipoles (R. H. Johnson et al., Electr. Letters, Vol. 22, pp. 591-593, 1986; R. H. Johnson et al., IEEE Trans. MTT, Vol. 35, pp. 1317-1321, 1987).
Distributed dipoles have been described with parallel slots in order to generate discrete parallel currents on proximal loopsides. They are provided with a flexible metallic backplane which allows the proximal loopsides to conform to cylindrical surfaces (R. H. Johnson et al., Proc. Hypert. Oncology 1988, T. Sugahara and M. Saito, eds., pp. 832-833, Taylor & Francis, 1989; A. W. Preece et al., Proc. 10th ESHO Symp., Amsterdam, p. 152, 1989).
Some improvements in SAR penetration and uniformity have been obtained with the use of lower frequency 27 MHz dipoles in a symmetric two side-to-side parallel dipole configuration (the Twin-Dipole). These lumped or distributed dipoles do exhibit some limitations in their use. They are low-efficiency devices due to both the spread of high intensity stray EM fields into open space in spite of the presence of the metallic backplane or box and to their short height, i.e., to a close distal loopside carrying the out-of-phase return current (IEEE Trans. Micr. Theory Techn., Vol. MTT-34, pp. 612-619, 1986).
The latter limitation has been removed by the implementation of a large (120 cm c.a.) ribbon-like conductive sheet applied against the body surface on which a high-intensity distributed current at 13.56 MHz is flowing, while the distal loopside is removed to a remote distance (80 cm c.a.) and does not contribute to the heating field. This device has been shown to be effective in penetration; however, its heating efficiency drops to the lowest level since the required radio frequency power of a few kilowatts is almost completely dispersed in the open space. Moreover, there is no practical way of controlling the local heating field distribution (H. Kato et al., J. Microw. Power, Vol. 18, pp. 331-336, 1983).
Improvements in heating penetration and uniformity with magnetic dipoles has been obtained with the development of a 27 MHz hybrid dipole applicator (Strahlentherapie, Vol. 9, p. 547, 1985). This operates by superimposing to the induced E-field of a twin-dipole device the unidirectional and coherent E-field generated by an auxiliary capacitive device. With this two different-element phased array applicator, the central SAR gradient typical of induced E-field loops disappears and a broad and deep SAR maximum appears by virtue of the positive interference of the two superimposed heating EM fields provided that the respective phase, amplitude and orientation are adjusted. Such an applicator can, however, be applied only to specific anatomic sites.